Optical imaging by fluorescence complements the various types of nuclear medical instruments, such as those using positron emission tomography (PET), gammatomography (SPECT, i.e. monophotonic emission tomography), X-ray imaging (e.g. digital radiography, X-ray tomography) and MRI (magnetic resonance imaging).
Optical imaging using fluorescence requires the injection into a human or animal organism of a marker (e.g. an antibody/fluorophore conjugate) that specifically targets an area of interest to the biologist or doctor, for example a malignant tumor of an organ. Below, “fluorophore” means any molecular or particulate structure able to emit light in response to luminous excitation (e.g. organic fluorophores, semiconductor nanocrystals, quantum boxes, etc.), and “marker” means the injected substance comprising such a fluorophore.
It is therefore possible to detect cancerous nodules by a technique that is much less invasive and destructive than ionizing radiation imaging (e.g. imaging using X-rays or radioactive tracers). Moreover, optical imaging systems offer good resolution at millimeter scale.
Finally, it should be noted that the equipment necessary for such optical imaging is relatively simple, comprising in particular a light source in the form of a compact laser diode, a detector in the form of a high-sensitivity camera, and motorized tables, and is of a very much lower cost than imaging equipment using ionizing radiation.
The simplest imaging systems using fluorescence include a light source (e.g. of fiber, laser, arc lamp, light-emitting diode type) and a filtered camera (filtered to avoid backscattering of the excitation light) for acquiring a fluorescent image; this is known as fluorescence reflectance imaging (FRI). The photons penetrate only a small distance (approximately 1 millimeter (mm) into the tissue and this technique can locate only markers on the surface of the tissue (e.g. only marked surface tumors in oncology).
If the marker (e.g. the marked tumor) is deep (for example 1 centimeter (cm) deep), it is impossible to locate it using FRI type acquisition alone, because of strong diffusion of the excitation photons and because of the photons emitted by the fluorophore. The light source is therefore moved, to “grid” the area to be analyzed and acquire as many images as possible of the light source. A complex reconstruction process based on all of the acquired images reconstructs the fluorescent image in “3D”: this is known as optical tomography. Small animals such as rats or mice can be imaged using a device of this kind.
When optically imaging the entire body of a small animal, it is generally possible either:                to acquire a global fluorescent image in two dimensions in a few seconds at most using the FRI technique, with the drawback of having no information as to depth in the tissue; or        to use the tomography technique, which produces an image in three dimensions but which requires, firstly, a series of images captured for respective different source/detector positions (acquisition time from 10 to 15 minutes for a field of approximately 1 cm2 and a step size of 2 mm) and, secondly, image reconstruction algorithms, with reconstruction time being proportional to the size of the area to be reconstructed. The difficulty encountered in “3D” reconstruction by tomography when a fluorophore is used therefore lies in the lack of information on the restricted area in which the markers are likely to be found (these markers are generally fixed to the organ to be observed, such as a malignant tumor, for example). It is then necessary to scan the whole of the object for imaging (for a mouse the total area is 50×75 mm2), with a step size of a few millimeters to avoid excessively long acquisition times, and information is lost between the excitation points.        
Moreover, image reconstruction algorithms usually start from a homogeneous distribution of sources of fluorescence in the entire body of the animal (i.e. the same, zero or non-zero, level of fluorescence). This therefore requires a large number of iterations before converging towards a reliable distribution of the sources of fluorescence and can therefore cause errors.
The excitation wavelength used is usually from 600 nanometer (nm) to 800 nm and the fluorophores (typically Cy5, Cy7, Alexa 633 or Alexa 750 cyanines) emit at wavelengths from 700 nm to 900 nm. In this range of wavelengths, autofluorescence of biological tissue is reduced compared to the bluegreen range (from 400 nm to 500 nm), but this undesirable autofluorescence is nevertheless always present (autofluorescence of biological tissue is caused by the presence of endogenous chromophores, such as porphyrines in haemoglobin, fluorescent proteins, etc.). Furthermore, with a fluorophore fixed to the organ to be detected, there is the drawback of the signal/noise ratio being relatively low.
A major drawback of those prior art optical imaging techniques using fluorescence is therefore essentially, with the FRI technique, the absence of depth information for locating the marker in the organ to be imaged and, with tomography, the difficulty of obtaining satisfactory depth (direction Z) information in addition to information relating to the surface of the organ (directions X and Y).
The paper “A dual fluorochrome probe for imaging proteases”, M. F. Kircher, R. Weissleder, L. Josephson, Bioconj. Chem. 2004, 15, 242 describes an optical imaging method using a dual marker formed of a nanoparticle functionalized by two organic fluorophores (Cy5.5 and Cy7 cyanines) to evaluate the activity of an enzyme in tissue. The ratio of the emissions from those two fluorophores is calculated to estimate the location in depth of the marked vector in the tissue. The marker is either excited at λ1=630 nm and analyzed at λ′1=700 nm for the Cy5.5 fluorophore, or excited at λ2=736 nm and analyzed at λ′2=800 nm for the Cy7 fluorophore, in accordance with the acquisition set-up illustrated in the appended FIG. 1, which shows that to each excitation wavelength λ1, λ2 there corresponds one and only one narrow spectral band around a wavelength λ′1 or λ′2 of maximum fluorescence
A major drawback of that optical imaging method using a dual marker is the complexity of fabrication of the marker, because of the operation of grafting the two fluorophores onto the same vector.
Another drawback of that method is that the relatively wide absorption/emission spectra of those two fluorophores overlap, which generates filtering difficulties for detecting each of the two fluorophores separately in order to be able to preserve the ratio of the two emissions yielding the estimate of the depth in the tissue. Moreover, the proximity of the absorption and emission wavelengths of the two markers generates spurious phenomena such as energy transfer and inhibition of fluorescence.
Another drawback of that method is that the emission spectrum of those two markers based on organic fluorophores can be offset because of interactions with the biological environment.
A final drawback of that dual marker method is that the use of the ratio of the intensity of the emissions of fluorescence of the two fluorophores to estimate the depth of the marker relies on two assumptions that are open to challenge, i.e. that the respective absorption coefficients a1 and a2 of the tissue at the two excitation wavelengths λ1 and λ2 are the same, as are the respective absorption coefficients a′1 and a′2 of the tissue at the emission wavelengths λ′1 and λ′2 (i.e. a1=a2 and a′1=a′2). These two assumptions are justified in the above publication by the mutual proximity firstly of the wavelengths λ1 and λ2, and secondly of the wavelengths λ′1 and λ′2. However, that proximity of the excitation and emission wavelengths of the two fluorophores used is found to penalize the optical filtering for measuring the fluorescence ratio, which degrades the quality with which the depth of the marker is evaluated in the tissue.
U.S. Pat. No. 5,370,119 describes a method of measuring the pH of an appropriate target that produces information as to the evolution of pH over time (its kinetic profile), comprising:                bringing the target to be analyzed into contact with a fluorescent marker having at least two excitation peaks and only one emission peak, with the emission spectrum being dependent on pH;        successively exciting the target treated in this way at the excitation wavelengths of the fluorescent marker; and        measuring the fluorescence emitted by the target at those excitation wavelengths and calculating the pH of the target from the ratio of the fluorescence signals emitted in response to the two excitation wavelengths by reading off the pH corresponding to the ratio obtained on a calibration curve for the marker, giving the ratio as a function of pH.        
Note that fluorescent markers of the organic fluorophore type such as fluoroscein used in the above document are not able to emit at a plurality of emission wavelengths in response to the plurality of excitation wavelengths, but at one and only one emission wavelength.
Note further that the measurement method of the above document does not provide an estimate of the three-dimensional location of the marker in the tissue or of the mean absorption coefficients of the tissue in relation to the excitation wavelengths.